1. Field of the Invention
The present invention is directed to a radiation converter of the type suitable for use in an x-ray system.
2. Description of the Prior Art
German OS 33 32 648 discloses a radiation converter embodied as an image intensifier. Such image intensifiers have an input window with a radiation absorber for generating light photons in a manner dependent on the radiation intensity of impinging radiation. Arranged downstream of the radiation absorber is a photocathode which generates electrons in a manner dependent on the light photons emerging from the radiation absorber. The electrons are accelerated onto an electron receiver by an electrode system. In the case of the image intensifier, the electron receiver is embodied as an output screen which generates light photons dependent on the impinging electrons.
U.S. Pat. No. 5,369,268 discloses an x-ray detector in which the photocathode is applied on a radiation absorber. The photocathode is arranged at a distance from and opposite an amorphous selenium layer of an output screen.
A further detector device is disclosed in German OS 44 29 925. In this case, a shadowmask produced from wires is provided on the radiation input side, a chevron plate being connected downstream of this shadowmask. In order to generate an image signal, a low-impedance anode structure is provided outside the detector on the rear side thereof. European Application 0 053 530 discloses a photodetector in which an electron multiplier and a detector anode are connected downstream of a photocathode in the radiation direction.
Since, in the case of medical examination of a patient, in contrast to nondestructive materials testing, the radiation loading must be kept as small as is technically practical, in order to minimize the radiation loading on the patient, efficient utilization of the radiation which penetrates through the patient and strikes the radiation receiver is the highest requirement. The smaller the radiation intensity on the radiation receiver, however, the smaller are the signals which can be derived from the radiation receiver. The distance between the signal levels and the noise signals likewise becomes smaller, which is associated with a poorer diagnosis capability of the image representations that can be generated on the basis of these signals. It is thus necessary to make a compromise between a small radiation loading on the patient and the radiation dose required for a good diagnosis capability of radiographic images of the patient that can be generated.
A photographic film is, for example, nothing more than a chemical amplifier which amplifies the ionization processes of the radiation in the microscopic region by many orders of magnitude and makes them visible in the macroscopic region.
Storage phosphor panels store the radiation shadow image of an object in latent fashion. By scanning the storage phosphor panel using a light beam, light photons are generated dependent on the latent image and are converted by a read-out with a photomultiplier into electrons which can be amplified virtually in noise-free fashion by up to a factor of 106, and converted into electrical signals. These electrical signals then are available for the image representation.
In x-ray image intensifiers, the geometrical size reduction which results due to the large input window and the smaller output window is used for intensifying the luminance, assistance for this being provided by the energy absorption of the electrons from the input fluorescent screen to the output fluorescent screen through an intervening acceleration field.
In the case of so-called flat panel image detectors, a layer which converts radiation into light and has CsI, for example, is brought into direct contact with a photodiode matrix made of amorphous silicon, so that the light photons generated by the layer due to incident radiation can be converted by means of the photodiode matrix into electrical signals, which are then available for the image representation. Since the light photons are not amplified by means of electrons, only relatively small signals can be derived from the photodiode matrix, which signals can be amplified only in a device connected downstream, e.g. an amplifier. Since the quantities of charge of these relatively small electrical signals then also must be additionally conducted, by means of complicated timing methods, from the large-area flat panel image detectors via relatively long lines as far as the amplifiers, the average noise, measured in electrons, is almost twice as great as the signal generated by individual x-ray quanta. Particularly for fluoroscopy, in which only small x-ray doses are applied, the signals which can be derived from the flat panel image detector are particularly small and near the region of the noise and thus require complicated artifact corrections. In fluoroscopy, as an example, the signals of every other beam scanning are used for correction purposes, so that nothing comparable to the customary image refresh rates can be achieved. Moreover, the dynamic range of the signals which can be derived from the flat panel image detector is greatly restricted.
In today's flat panel image detectors, predominantly a-Si:H read-out panels are used as electron detectors. Operation of such flat panel image detectors in different operating modes, such as fluoroscopy and radiography, which differ by dose factors of 100–1000, requires a high computation complexity. In the transition from the radiography operating mode, operated with a high dose, to the fluoroscopy operating mode, operated with a low dose, residual images in the a-Si:H read-out panel must be removed computationally by subtraction.